Examination of a biological tissue using photon migration between a plurality of input and detection locations

ABSTRACT

A spectroscopic method and system for examination of biological tissue includes multiple input ports optically connected to at least one light source, multiple detection ports optically connected to at least one detector, a radiation pattern controller coupled to the light source and detector, and a processor. The multiple input ports are arranged to introduce light at input locations into biological tissue and the multiple detection ports are arranged to collect light from detection locations of the biological tissue. The radiation pattern controller is constructed to control patterns of light introduced from the multiple input ports and constructed to control detection of light migrating to the multiple detection ports. The processor is operatively connected to the radiation pattern controller and connected to receive detector signals from the detector, and is constructed to examine a tissue region based on the introduced and detected light patterns.

This application is a Continuation of U.S. Ser. No. 08/356,162, filedDec. 16, 1994, now U.S. Pat. No. 5,807,263, which was filed under 35 USC§371 of PCT/US93/05868, filed Jun. 17, 1993, which is acontinuation-in-part of U.S. Ser. No. 07/900,197, filed Jun. 17, 1992,now U.S. Pat. No. 5,353,789.

BACKGROUND OF THE INVENTION

This invention relates to examination and imaging of biological tissueusing visible or infra-red radiation.

Traditionally, potentially harmful ionizing radiation (for example,X-ray or λ-ray) has been used to image biological tissue. This radiationpropagates in the tissue on straight, ballistic tracks, i.e., scatteringof the radiation is negligible. Thus, imaging is based on evaluation ofthe absorption levels of different tissue types. For example, inroentgenography the X-ray film contains darker and lighter spots. Inmore complicated systems, such as computerized tomography (CT), across-sectional picture of human organs is created by transmitting X-rayradiation through a section of the human body at different angles and byelectronically detecting the variation in X-ray transmission. Thedetected intensity information is digitally stored in a computer whichreconstructs the X-ray absorption of the tissue at a multiplicity ofpoints located in one cross-sectional plane.

Near infra-red radiation (NIR) has been used to study non-invasively theoxygen metabolism in tissue (for example, the brain, finger, or earlobe). Using visible, NIR and infra-red (IR) radiation for medicalimaging could bring several advantages. In the NIR or IR range thecontrast factor between a tumor and a tissue is much larger than in theX-ray range. In addition, the visible to IR radiation is preferred overthe X-ray radiation since it is non-ionizing; thus, it potentiallycauses fewer side effects. However, with lower energy radiation, such asvisible or infra-red radiation, the radiation is strongly scattered andabsorbed in biological tissue, and the migration path cannot beapproximated by a straight line, making inapplicable certain aspects ofcross-sectional imaging techniques.

Recently, certain approaches to NIR imaging have been suggested. Oneapproach undertaken by Oda et al. in “Non-Invasive HemoglobinOxygenation Monitor and Computerized Tomography of NIR Spectrometry,”SPIE Vol. 1431, p. 284, 1991, utilizes NIR radiation in an analogous wayto the use of X-ray radiation in an X-ray CT. In this device, the X-raysource is replaced by three laser diodes emitting light in the NIRrange. The NIR-CT uses a set of photomultipliers to detect the light ofthe three laser diodes transmitted through the imaged tissue. Thedetected data are manipulated by a computer of the original X-ray CTscanner system in the same way as the detected X-ray data would be.

Different approaches were suggested by S. R. Arriadge et al. in“Reconstruction Methods for Infra-red Absorption Imaging,” SPIE Vol.1431, p. 204, 1991; F. A. Grünbaum et al. in “Diffuse Tomography,” SPIEVol. 1431, p. 232, 1991; B. Chance et al., SPIE Vol. 1431 (1991), p. 84,p. 180, and p. 264; and others who recognized the scattering aspect ofthe non-ionizing radiation and its importance in imaging. None of thosetechniques have fully satisfied all situations.

In summary, there continues to be a need for an improved imaging systemwhich utilizes visible or IR radiation of wavelengths sensitive toendogenous or exogenous pigments.

SUMMARY OF THE INVENTION

The invention relates to systems and methods for spectroscopicexamination of a subject positioned between input and detection ports ofthe spectroscopic system applied to the subject.

According to one aspect of the invention, a spectroscopic systemincludes at least one light source adapted to introduce, at multipleinput ports, electromagnetic non-ionizing radiation of a knowntime-varying pattern of photon density and of a wavelength selected tobe scattered and absorbed while migrating in the subject, the inputports being placed at selected locations on the subject to probe aselected quality of the subject; and radiation pattern control meansadapted to achieve selected a time relationship of the introducedpatterns to form resulting radiation that possesses a substantialgradient in photon density as a result of the interaction of theintroduced patterns emanating from the input ports, the radiation beingscattered and absorbed in migration paths in the subject. The gradientin photon density may be achieved by encoding the introduced radiationpatterns with a selected difference in their relative amplitude,relative phase, relative frequency or relative time. The system alsoincludes a detector adapted to detect over time, at a detection portplaced at a selected location on the subject, the radiation that hasmigrated in the subject; processing means adapted to process signals ofthe detected radiation in relation to the introduced radiation to createprocessed data indicative of the influence of the subject upon thegradient of photon density; and evaluation means adapted to examine thesubject by correlating the processed data with the locations of theinput and output ports.

Preferred embodiments of this aspect of the invention includedisplacement means adapted to move synchronously all the optical inputports or move the detection ports to another location on a predeterminedgeometric pattern; at this location the examination of the subject isperformed.

According to another aspect of the invention, a spectroscopic systemincludes at least one light source adapted to introduce, at multipleinput ports, electromagnetic non-ionizing radiation of a knowntime-varying pattern of photon density and of a wavelength selected tobe scattered and absorbed while migrating in the subject, the inputports being placed at selected locations on the subject to probe aselected quality of the subject; radiation pattern control means adaptedto achieve a selected time relationship of the introduced patterns toform resulting radiation that possesses a substantial gradient in photondensity as a result of the interaction of the introduced patternsemanating from the input ports, the radiation being scattered andabsorbed in migration paths in the subject. The system also includes adetector adapted to detect over time, at a detection port placed at aselected location on the subject, the radiation that has migrated in thesubject; displacement means adapted to move the detection port tovarious locations on a predetermined geometric pattern, the variouslocations being used to detect over time radiation that has migrated inthe subject; processing means adapted to process signals of the detectedradiation in relation to the introduced radiation to create processeddata indicative of the influence of the subject upon the gradient ofphoton density; and evaluation means adapted to examine the subject bycorrelating the processed data with the locations of the input andoutput ports.

According to another aspect of the invention, a spectroscopic systemincludes at least one light source adapted to introduce, at multipleinput ports, electromagnetic non-ionizing radiation of a knowntime-varying pattern of photon density and of a wavelength selected tobe scattered and absorbed while migrating in the subject, the inputports being placed at selected locations on the subject to probe aselected quality of the subject; radiation pattern control means adaptedto achieve a selected time relationship of the introduced patterns toform resulting radiation that possesses a substantial gradient in photondensity as a result of the interaction of the introduced patternsemanating from the input ports, the radiation being scattered andabsorbed in migration paths in the subject. The system also includes atleast one detector adapted to detect over time, at multiple detectionports placed at selected locations on the subject, the radiation thathas migrated in the subject; processing means adapted to process signalsof the detected radiation in relation to the introduced radiation tocreate processed data indicative of the influence of the subject uponthe gradient of photon density, and evaluation means adapted to examinethe subject by correlating the processed data with the locations of theinput and output ports.

Preferred embodiments of this aspect of the invention includedisplacement means adapted to move at least one of the detection portsto another location on a predetermined geometric pattern, the otherlocation being used to perform the examination of the subject.

Preferred embodiments of this aspect of the invention include rotationmeans adapted to rotate synchronously the optical input ports whileintroducing the resulting radiation along a predetermined geometricpattern, the input port rotation being used to perform the examinationof a region of the subject.

Preferred embodiments of the above described aspects of the inventionare also used to locate a fluorescent constituent of interest in thesubject; the wavelength of the introduced radiation is selected to beabsorbed in the fluorescent constituent, the detected radiation isemitted from the fluorescent constituent and processed to determinelocation of the fluorescent constituent.

According to another aspect of the invention, a spectroscopic systemincludes a light source adapted to introduce, at an input port,electromagnetic non-ionizing radiation of a known time-varying patternof photon density and of a wavelength selected to be scattered andabsorbed while migrating in the subject, the input port being placed ata selected location on the subject to probe a selected quality of thesubject; detectors adapted to detect over time, at multiple detectionports placed at selected locations on the subject, the radiation thathas migrated in the subject; the time relationship of the detection overtime, at the detection ports, being selected to observe a gradient inphoton density formed as a result of the interaction of the introducedradiation with the subject. The system also includes processing meansadapted to process signals of the detected radiation in relation to theintroduced radiation to create processed data indicative of theinfluence of the subject upon the gradient of photon density, andevaluation means adapted to examine the subject by correlating theprocessed data with the locations of the input and output ports.

Preferred embodiments of this aspect of the invention of the inventioninclude displacement means adapted to move at least one of the detectionports to another location on a predetermined geometric pattern, theother location being used to perform the examination of the subject.

According to another aspect of the invention, a spectroscopic systemincludes a light source adapted to introduce, at an input port,electromagnetic non-ionizing radiation of a known time-varying patternof photon density and of a wavelength selected to be scattered andabsorbed by a fluorescent constituent while migrating in the subject,the input port being placed at a selected location on the subject tolocate the fluorescent constituent of the subject; detectors adapted todetect over time, at multiple detection parts placed at selectedlocations on the subject, fluorescent radiation that has migrated in thesubject. The system also includes processing means adapted to processsignals of the detected radiation in relation to the introducedradiation to create processed data indicative of location of thefluorescent constituent of the subject, and evaluation means adapted toexamine the subject by correlating the processed data with the locationsof the input and output ports.

Preferred embodiments of this aspect of the invention includedisplacement means adapted to move at least one of the detection portsto another location on a predetermined geometric pattern, the otherlocation being used to locate the fluorescent constituent of thesubject.

Preferred embodiments of the above-described aspects of the inventionuse one or more of the following features:

The time-varying pattern comprises radiation of a selected wavelengthintensity modulated at a selected frequency. The radiation patterncontrol means are further adapted to control a selected phaserelationship between the modulated radiation patterns introduced fromeach of the input ports having to produce in at least one direction asteep phase change and a sharp minimum in the intensity of theradiation.

The radiation pattern control means are further adapted to impose on allthe introduced radiation patterns an identical time-varying phasecomponent thereby changing the spatial orientation of the direction ofthe steep phase change and the sharp minimum in the intensity of theradiation.

The time-varying pattern comprises radiation of a selected wavelengthintensity modulated at a selected frequency. The radiation patterncontrol means are further adapted to control a selected frequencyrelationship between the modulated radiation patterns introduced fromeach of the input ports having to produce in at least one direction asteep phase change and a sharp minimum in the intensity of theradiation.

The time-varying pattern comprises radiation of a selected wavelengthintensity modulated at a selected frequency. The radiation patterncontrol means are further adapted to control a selected amplituderelationship between the modulated radiation patterns introduced fromeach of the input ports having to produce in at least one direction asteep phase change and a sharp minimum in the intensity of theradiation.

The radiation pattern control means are further adapted to add to allthe introduced radiation patterns an identical time-varying amplitudecomponent thereby changing the spatial orientation of the direction ofthe steep phase change and the sharp minimum in the intensity of theradiation.

The radiation is modulated at a frequency that enables resolution of thephase shift that originates during migration of photons in the subject.

The frequency is on the order of 10⁸ Hz.

The processing means further adapted to determine the phase or theintensity of the radiation altered by scattering and absorption in thesubject.

The wavelength of the radiation is susceptible to changes in anendogenous or exogenous tissue pigment of the subject.

The gradient in photon density may also be achieved by encoding theintroduced radiation patterns with a selected difference in theirrelative amplitude, relative phase, relative frequency or relative time.

Other advantages and features of the invention will be apparent from thefollowing description of the preferred embodiment and from the claims.

BRIEF DESCRIPTION OF THE DRAWING

FIGS. 1, 1A and 1B show diagrammatically phase modulation imagingsystems employing several input ports and one detection port inaccordance with the present invention.

FIG. 2 is a block diagram of the phase modulation imaging systemincluding several input ports and several detection ports in accordancewith the present invention.

FIG. 2A depicts a phased array transmitter that radiates a directionalbeam.

FIG. 2B depicts sequencing of the phases of an antiphase multi-elementarray to achieve an electronic scan of the photon density gradient inaccordance with the present invention.

FIG. 2C depicts four element antiphased array designed for a conicalscan of the photon density gradient in accordance with the presentinvention.

FIG. 2D depicts the input and output port arrangement of an imagingsystem in accordance with the present invention.

FIGS. 3 and 3A depict an imaging system for detection of a hiddenfluorescing object in accordance with the present invention.

FIG. 4 is a block diagram of an alternative embodiment of a dualwavelength PMS system.

FIG. 4A is a schematic diagram of an oscillator circuit of FIG. 4.

FIG. 4B is a schematic diagram of a PMT heterodyne modulation and mixingnetwork shown in FIG. 4.

FIG. 4C is a schematic diagram of an AGC circuit shown in FIG. 4.

FIG. 4D is a schematic diagram of a phase detector circuit shown in FIG.4.

FIGS. 5A, 5B, and 5C illustrate changes in optical field propagating ina strongly scattering medium which includes a strongly absorbingcomponent.

FIG. 6 shows an experimental arrangement of a two element phased arrayused in an interference experiment.

FIGS. 6A, 6B, and 6C show detected interference patterns of twodiffusive waves.

FIG. 7 displays the phase shifts measured for a two element array (curveA), and for a single source (curve B).

FIG. 8A depicts an experimental arrangement of sources of a four elementphased array and a detector.

FIGS. 8B and 8C display the intensities and the phase shifts measuredfor the four element array of FIG. 8A, respectively.

FIG. 9A depicts an experimental arrangement of sources of a four elementphased array, a detector, and a strongly absorbing object.

FIGS. 9B, 9C display respectively the intensities and the phase shiftsmeasured for the four element array of FIG. 9A scanning absorbingobjects of different sizes.

FIG. 9D displays the phase shifts measured for the four element array ofFIG. 9A scanning absorbing objects of different absorption coefficients.

FIG. 10A an experimental arrangement of sources of a four element phasedarray, a detector, and two strongly absorbing objects.

FIG. 10B displays the phase shifts measured for the four element arrayof FIG. 10A scanning two absorbing objects of different sizes.

FIG. 11 depict diagrammatically a single wavelength localization systemutilizing a conical scanner.

FIGS. 11A and 11B depict diagrammatically imaging systems utilizing oneor two dimensional phased array transmitters.

FIGS. 12A and 12B depict an imaging system comprising a two dimensionalphased array transmitter and detection array.

DESCRIPTION OF THE PREFERRED EMBODIMENT

Imaging system embodiments of the present invention based uponinterference effects of radiation migrating in a subject havingscattering and absorptive properties are shown in FIGS. 1, 2, and 3. Thesystems effectively utilize, in this scattering medium, a directionalbeam of visible or IR radiation generated and/or detected by an array ofsources and/or detectors, respectively. For instance, in the case of anarray of sources, each source is placed at a selected location in thearray and emits intensity modulated radiation, preferably coherentradiation from a laser diode, of a selected intensity and phase. Thecriteria for selecting the source locations, the intensities, and thephases of the respective sources is the shape of the desired beam thatat any time point possesses a substantial photon density gradientproduced by interference effects of radiation from the various sources.This gradient of photon density is localized and has directionalproperties. Overall, the resulting radiation formed by interference ofthe radiation of the individual sources migrates in a selected directionin the subject. In an antiphase system, the wavefront of the beam hassections of equal photon density separated by a sharp localized changein photon density. Selected different locations of the photon densitygradient are shown in FIG. 2B.

In general, the wavefront propagates in the selected direction in thesubject and the gradient of photon density is localized in one or moreplanes extending from the source array in a selected direction. If thesubject includes a localized object having different scattering andabsorptive properties from those of the surrounding environment, thepropagating radiated field is perturbed. This perturbation is detectedand from the source detector geometry the perturbing object can belocated.

Referring to the embodiment of FIGS. 1 and 1A, the imaging systemutilizes an array of laser diodes 12, 14, 16, and 18 for introducinglight into the tissue at selected locations. The geometry of opticalinput ports 11, 13, 15, 17 and of an optical output port 19 is selectedto examine a specific part of the tissue. From the known geometry of theoptical input ports and the detection port and from the shape of theintroduced and detected radiation, a computer can locate a hidden object9 of examined tissue 8 (for example, a head or breast). A masteroscillator 22, which operates at 200 MHz, excites laser diodes 12through 18, that emit light of a selected wavelength (e.g., 760 nm). Thelight from each laser diode is conducted to the respective input portplaced on a subject via a set optical fibers. A detector 24 detects thelight that has migrated through the examined tissue. Preferably,detector 24 includes a photomultiplier tube (e.g., Hamamatsu R928)powered by a high voltage supply which outputs about 900 V in order toensure a high gain. A local oscillator 26 operating at a convenientoffset frequency (e.g., 25 kHz) sends a signal to a mixer 28 and areference signal to detector 24. Accordingly, an output waveform 25 fromdetector 24 is at a carrier frequency equal to the difference of thedetected and reference frequency, i.e., 25 kHz.

Detector 24 (for example, PMT Hamamatsu R928 or Hamamatsu R1645u)detects the scattered and absorbed light that has migrated through thesubject. Detection port 19 is located several centimeters from thelocation of the input ports. The PMT detector is connected to thesubject by the fiber optic guide, or, alternatively, may be directlyplaced on the subject. It has been found that the most cost-effectivedetector for measuring signals of frequencies on the order of 10⁸ Hz isHamamatsu R928. However, the Hamamatsu R1645u detector is preferred dueto its high precision. The second dynode of the PMT of detector 24 ismodulated by 200.025 MHz signal 27 so that the 25 kHz hetrodyned signal25 is received by a phase detector 30. Phase detector 30 also receivesreference signal 29 from mixer 28. If phase detector 30 is a lock-inamplifier then the output signals are the phase shift and the intensityof the detected signal. Both the phase shift and the intensity of thedetected light characterize the migration path of photons in the subject(e.g., the brain tissue).

Alternatively, a tunable dye laser or other laser source connected to awide band acousto-optical modulator operating at the carrier frequency,e.g., 200 MHz can be used instead of the laser diode. Theacousto-optical modulator modulates the intensity of the light emittedby the laser at the selected carrier frequency.

The invention also envisions using only one source of coherent lightthat irradiates one end of several optical fibers at the same time. Theother end of each fiber is placed on the subject at a selected inputport location. This source radiates light of a selected time varyingpattern. The phase relationship and the intensity of the light carriedby each fiber is varied by creating a time delay (e.g., different fiberlength) and by coupling different amounts of light into each fiber.

FIG. 1B shows diagrammatically an imaging system of FIG. 1 furtheradapted to encode the transmitted light using an offset frequency.Oscillators 22 a, 22 b, 22 c and 22 d drive four laser diodes atfrequencies 30.025 MHz, 30.035 MHz, 30.045 MHz and 30.055 MHz,respectively. The laser diodes introduce the light that migrates intissue 8 and is collected at detection port 19 and detected by PMTdetector 24. Local oscillator 26 provides a 30 MHz reference signal todetector 24 that outputs a detection signal having 25 kHz, 35 kHz, 45kHz and 55 kHz frequency components. Each component signal is phasedetected at a corresponding phase detector (30 a, 30 b, 30 c and 30 d)having a suitable frequency filter. The phase detectors provide a phaseshift, migration pathlength and amplitude for each frequency.

The imaging systems of FIGS. 1, 2, and 3 are shown to have a lightsource of a single wavelength; however, a dual wavelength imaging systemis also envisioned according to this invention. In the dual wavelengthimaging system two laser diodes or a tunable wavelength laser generatelight of two wavelengths that is coupled to an optical fiber. Such asystem will now be described.

A dual wavelength operation is shown in FIG. 4. The system includes amaster oscillator 60 operating at 200 MHz and an oscillator 62 operatingat 200.025 MHz which is offset 25 kHz from the master oscillatorfrequency. The offset frequency of 25 kHz is a convenient frequency forphase detection in this system; however, other offset frequencies ashigh as a few megahertz can be used. Oscillator 60 alternatively drivestwo sets of laser diodes 64 a, 64 b, . . . , 64 n and 66 a, 66 b, . . ., 66 n using switches 61 a, 61 b, . . . , 66 n. These switches aredriven electronically to couple a selected wavelength into the opticalfiber and also to achieve a selected radiation pattern resulting fromthe radiation emanating from the individual fibers. An output 8 mm fibercoupler 72 collects photons for an R928 PMT detector 74. The seconddynode (shown in FIG. 3B) of PMT 74 is modulated with a 200.025 MHzreference signal generated by oscillator 62 and amplified by anamplifier 63. Thus, the output signal of the PMT detector has afrequency of 25 kHz. PMT detector 74 alternately detects light of thetwo laser diodes that has migrated in the tissue and producescorresponding output signals, which are filtered by a filter 76 andleveled by an automatic gain control (AGC) circuit 79. A referencesignal of 25 kHz is produced in a mixer 65 by mixing the 200 and 200.025MHz oscillator signals. The reference 25 kHz signal is also leveledusing the second AGC 77 and fed into a phase detector 79. Phase detector79 generates a signal indicative of the phase of each output signalrelative to the phase of the reference signal. The outputs of phasedetector 79 are alternately selected by an electronic switch 80,filtered, and then input to an adder 82 and a subtractor 81 to producesum and difference signals proportional to <L>_(λ1)+<L>_(λ2) and<L>_(λ1)−<L>_(λ2). The difference and sum signals are then used tocalculate changes in the probed pigment and in the blood volume,respectively.

A schematic diagram of preferred oscillator 60 or 62 is shown in FIG.4A. This circuit has a drift of only 0.03 degrees/hr. (Weng, et al.,“Measurement of Biological Tissue Metabolism Using Phase ModulationSpectroscopic Measurement,” SPIE, Vol. 143, p. 161, 1991, which isincorporated herein by reference). The crystal is neutralized, whichenables operation at resonance, and thus achieves long-term stability.The respective crystals of oscillators 60 and 62 are offset from eachother by 25 kHz. This circuit provides a sufficient output to directlydrive a 5 mW laser diode.

A modulation circuit 75 for the second dynode of the PMT is shown inFIG. 4B. This circuit uses a resonant circuit 75 a with an impedance of20,000 ohms instead of the usual 50Ω load with very high powerdissipation, providing a 50 V drive of the photomultiplier dynode whiledissipating only a few watts of power.

To obtain stable operation of the phase detector, a stable input signalis required. The 25 kHz AGC circuit 77, 78 illustrated in FIG. 4Cincludes an MC 1350 integrated circuit Ul, featuring wide range AGC foruse as an amplifier. The signal amplitude is controlled by a feedbacknetwork, as shown. A major reason for the accurate detection of phasechanges by the PMT system is that the phase detector input signal levelis kept nearly constant by the AGC circuit. Since the input voltagechange of between 2 and 6 volts causes variation in the phase shift ofonly 0.2%, the AGC circuit eliminates the need for a very stable highvoltage power supply.

A preferred phase detector circuit is shown in FIG. 4D. Two sinusoidalsignals (the measurement signal and the reference signal) aretransformed to a square wave signal by a Schmitt trigger circuit 79 a.The phase of the square wave signal is shifted by an RC change (composedof R11, R12, C8), which makes it possible to change the measuring range.The detector further includes a 74HC221 integrated circuit. The lock-inamplifier technique obtained to derive the difference of the phase andamplitude of the two signals has the highest signal to noise ratiopossible for this type of equipment.

The above-described systems utilize the carrier frequency on the orderof 10⁸ Hz which is sufficiently fast to resolve the phase shift of thedetected light. The characteristic time, the time it takes for a photonto migrate between an input port and an output port, is severalnanoseconds. The sensitivity of the system is high, approximately 70°per nanosecond or 3° per centimeter change of pathlength, as observed inexperimental models. Selection of the modulation frequency also dependson the desired penetration depth and resolution of the imaging systemthat will be described below. If deep penetration is desired, a lowmodulation frequency (e.g., 40 MHz) is selected, and if shallowpenetration is needed, modulation frequencies as high as 10⁹ Hz can beused.

Referring to FIGS. 1 and 1A, a master oscillator 22 operates at amodulation frequency in the range of 40 to 400 MHz selected according tothe desired penetration depth of the optical field. The array of laserdiodes 12, 14, 16, and 18 generates a highly directional radiationpattern, which is employed in the tissue examination.

In one preferred mode of operation, laser diodes 12 to 18 operate in aphased array pattern which is introduced into the tissue and detected bya single PMT detector 30. Master oscillator 22 operating at 200 MHzdrives a multi-channel phased splitter which gives outputs atpredetermined phases. Input ports 11 through 17 are located at selecteddistances and an appropriate phasing of the array creates a directionalbeam and enables scanning of the optical field in two dimensions acrossthe tissue, as shown in FIGS. 2A, 2B, and 2D. After migrating throughthe tissue, the optical field is collected in a large area fiber onselected locations 19. The detected signals are heterodyned in the PMTdetector 24 by utilizing the output of local oscillator 26, operating ata 25 kHz offset frequency, to detector 24. The resulting 25 kHz signalis phase detected with respect to the output signal 29 of mixer 28 anddetector 24. Phase detector 30 outputs the phase and the intensity ofsignal 25. The detected phase shifts and intensities are stored and usedfor construction of an image of the subject. This is performed bycomputer control 34, which governs the operation of the system.

FIG. 2 depicts a phase modulation imaging system comprising an inputport array for introducing radiation and a detection port array fordetecting radiation that has migrated in the subject. The operation ofthe system is controlled by computer control 34, which coordinates atransmitter unit 32 with a receiver unit 42. Transmitter unit 32comprises several sources of visible or IR radiation adapted tointroduce a selected time-varying pattern of photon density into subject8 by array of input ports 31, 33, 35, and 37. Receiver unit 42 detectsradiation that has migrated in the subject from the input port array toan array of detectors 39, 41, 42, and 47.

The radiation sources of transmitter unit 32 are intensity modulated ata frequency in the range of 40 MHz to 200 MHz, as described for theimaging system of FIG. 1. Receiver unit 42 detects and processes theradiation using the same principles of the phase and amplitude detectionas described above. The signal detected at individual ports can bephased using appropriate delays.

Several modes of operation of the transmitter array and receiver arrayare described in FIGS. 2A, 2B, 2C, and 2D. Referring to FIG. 2A, it hasbeen known, that for a simple horizontal linear array of N identicalelements radiating amplitude modulated light spaced a distance, d,apart. The radiating wavefront is created by the interference effect. Ifall elements radiate in phase the wavefront propagates in a directionperpendicular to the array. However, by appropriately phasing theradiating elements, the resulting beam can scan space in two dimensions.We consider the phases of the signal along the plane A—A whose normalmakes an angle θ₀ with respect to the array normal. The phase of thesignal from the first radiator lags the phase of the second radiator bya phase angle (2π/λ)d sin θ₀ because the signal from the second radiatorhas to travel a distance d sin θ₀ longer than the signal from the firstradiator to reach plane A—A. Similarly, the phase of the signal from then^(th) radiator leads that from the first radiator by an angle n(2π/λ))dsin θ₀. Thus, the signals from the various radiators can be adjusted tobe in-phase along the A—A plane, if the phase of each radiator isincreased by (2π/λ)d sin θ₀. Consequently, at a point on the wavefrontin the far field of the transmitter array the signals from the Nradiators will add up in phase, i.e., the intensity of the totalnormalized signal is a sum of the signals from the individual sources.The constructed pattern has a well defined directional characteristicand a well pronounced angular dependence, i.e., the transmitter patternhas a well defined transfer characteristic of the transmitter withrespect to the angle θ₀.

FIG. 2B depicts an arrangement of phases for the sources the system ofFIG. 2 operating in one preferred mode of operation. The array of fivesources is divided into two or more portions that are phased 180° apart.Each portion has at least one source. The sources of each portionradiate amplitude modulated light of equal intensity and are spaced sothat the resulting beam of two or more equally phased sources has asubstantially flat wavefront, i.e., no gradient of photon density. Onthe other hand, there is a sharp 180° phase transition, a large gradientin photon density between two antiphased portions of the array. Thus,the radiated field possesses an amplitude null and a phase transition of180° (i.e. crossover phase), which is due to the large gradient ofphoton density.

Electronic scanning is performed by appropriately varying theapportionment of 0° and 180° phases on the sources. The five elementarray of FIG. 2B can have the 180° phase transition along four differentparallel planes extending from the array. Scanning is achieved byelectronically switching the sources by 180°, so that the photon densitygradient moves in the direction parallel to the location of the sources.

Using the principles described in FIGS. 2A and 2B, a conical scan of adirectional beam possessing at least one substantial photon densitygradient can be accomplished using a four element antiphased array, asshown in FIG. 2C. The laser diodes are antiphased using a push pulltransformer. The phasing and amplitude of four laser diodes S₁, S₂, S₃,and S₄ arranged into a two dimensional array is modified sequentiallyusing the switches Sw₁, Sw₂, Sw₃, and Sw₆ and inductances L₁, L₂, L₃,and L₄.

FIG. 2D shows a possible arrangement of the transmitter array and thereceiver array. The above described directional beam enters subject 8 atthe transmitter array location and is pointed to hidden absorber 9 whichperturbs the migrating beam. The field perturbation is measured by thereceiver array. Scanning of the transmitter array or the receiver arrayis envisioned by the present invention.

A hidden absorber that includes a fluorescent constituent is detectedusing a selected excitation wavelength of the laser sources of thetransmitter array. Then, the radiation is absorbed, and almost instantlya fluorescent radiation of a different wavelength is re-emitted. There-emitted radiation propagating in all directions is detected by thereceiver array.

FIG. 3 depicts a phase modulation imaging system comprising one inputport and several arrays of detection ports. This system operatescomparably to the systems of FIGS. 1 and 2. The 754 nm light of a laserdiode 48 is amplitude modulated using master oscillator 22. The light iscoupled to subject 8 using an input port 49. The amplitude modulatedlight migrates in the subject and is scattered from hidden object 9. Itis also expected that hidden object 9 has a different effective index ofrefraction than subject 8. The migrating radiation is governed by thelaws of diffusional wave optics that are described below. The scatteredradiation migrates in several directions and is detected by detectionsystems 50, 52, and 54.

Ports 51, 53, and 55 of the detection systems can include either largearea fibers or arrays of detection ports. If large area fibers are usedthen detector systems 50, 52, and 54 correspond to detector 24 of FIG.1. If arrays detection ports are used, then each of detector systems 50,52, and 54 includes several individual PMT detectors. The PMT detectorsof each detector system are phased utilizing a selected phase mode, asdescribed above. The phasing is controlled by the computer control. Thedetected signals are heterodyned at the PMT detectors and sent to aphase detector 58. Phase detector 58 detects alternatively theheterodyned signals using a switch 56. Operation of phase detector 58 issimilar to the operation of phase detector 30 of FIG. 1. The detectedphase and amplitude are alternatively sent to the computer control usinga switch 56 a. Even thought only one phase detector is shown in FIG. 3,the invention envisions use of several phase detectors.

If hidden absorber 9 includes a fluorescent constituent, laser diode 48is selected to introduce an excitation wavelength (e.g., 754 nm). Theintroduced, intensity modulated radiation, excites the fluorescentconstituent which re-emits radiation in all directions, as shown in FIG.3. The re-emitted radiation is detected using detector systems 50, 52,and 54. To increase the system resolution, each detector can befurnished with an interference filter selected to pass only thefluorescent radiation.

FIG. 3A shows diagrammatically an imaging system used for detection of afluorescing object 9. The system is a modified version of the system ofFIG. 3 wherein a four element phase array 47 introduces a 200 MHz lightof a 0° and 180° phase. The diffusion wave emitted from array 47 isre-emitted by object 9 and detected by ports 51, 53 and 55 and processedas described in connection with FIG. 3. Array 47 effectively codes theillumination light. Thus, when array 47 is rotated about the examinedorgan with object 9, the receivers will contain informationcorresponding to the orientation of the object. Each detection port alsoincludes a filter that passes only the fluorescent radiation; thisimproves the resolution of the system.

The interference of several waves, as described in FIG. 2A, has beenlong known in a non-scattering medium, wherein the radiation propagateson a straight line, but not in a strongly scattering medium. Referringto FIGS. 6, 6A, 6B, and 6C, in a simple experiment, interference of twodifferent diffusive waves in a strongly scattering medium wasdemonstrated. Propagation of visible IR radiation in a scattering mediumsuch as tissue can be described by diffusion of photons, and thus wedescribe it as a diffusive wave that exhibit refraction, diffraction andinterference. The diffusive waves, which can be visualized as “ripplesof brightness,” represent a scalar, over-damped traveling waves of lightenergy density.

Referring to FIG. 6, the two laser diodes were separated at a distanceof 4 cm and 1.2 cm from the detection port. The intensity modulatedlight of the two laser diodes at frequency 200 MHz was sent through twooptical fibers to a container with an Intralipid™ suspension. The sourcedetector distance was varied by moving the optical port of the detectionfiber along a line parallel to the position of the sources. FIGS. 6A,6B, and 6C show measured maxima and minima of the optical fieldmigrating in the medium. This data demonstrates interference between twodiffusive waves created by two coherent emitting sources of phasedifference 180 degrees. FIG. 7 summarizes the experiment, wherein thedisplacement of the detector is plotted against the phase shift measuredby the detector. The phase shift displays the steepest part of thetrace, curve A, (slope of about 360°/cm) at the displacement of about2.25 cm. Curve B is measured with an optical field of source S₂. Here,the measured slope is about 30°/cm. When comparing curves A and B wedemonstrate much higher sensitivity of the null detection of the twoelement array contrasted with a diminished sensitivity to the detectordisplacement when using a single source arrangement. The sensitivity ofthe two source arrangement is increased by about a factor of 10. Thesensitivity is further increased when using four or more element phasedarray, which sharpens the photon density gradient and thus provides ahigher resolution for locating the hidden object.

In a strongly scattering medium, the emitted photons undergo a largenumber of collisions and their migration can be determined by applyingthe diffusion equation. The diffusion equation for photons in auniformly scattering medium was solved by E. Gratton et al., “Thepossibility of a near infrared optical imaging system using frequencydomain methods.” in Mind Brian Imaging Program, Japan 1990; and by J.Fishkin et al., “Diffusion of intensity modulated near-infrared light inturbid media”, SPIE Vol. 1413 (1991) p. 122. A solution of the diffusionequation was obtained for the light of a point source (at r=0) radiatingS{1+M exp[−i(ωt+e)]} photons, wherein S is the source strength(photons/sec.), M is the modulation of the source at frequency ω, and eis an arbitrary phase. The photon intensity can be calculated as

I(r,t)=c*ρ(r,t)

wherein ρ(r,t) is the photon density and c=10⁸ m/s is the velocity oflight.

When solving the diffusion equation using a spherical-harmonicsapproximation in a non-absorbing medium for the density of photonsρ(r,t) than

I(r,t)=(I ₀ /Dr)+(I ₀ /Dr)exp[−r(ω/2cD)^(½)]×exp[ir(ω/2cD)^(½)−i(ωt+e)],

wherein the diffusion constant D is ⅓ of the mean free path. In theabsence of an amplitude modulated signal (ω=0) the solution correspondsto a spherical wave propagating without attenuation. For a non-zerofrequency, the amplitude of the signal at a frequency ω decreasesexponentially. The light wave front the emitted advances at the constantvelocity V

 V=(2Dcω)^(½)

and has wavelength

λ=2π(2cD/ω)^(½)

The above equations show that higher modulation frequencies yieldshorter effective wavelengths, and smaller diffusion constants also giveshorter effective wavelengths. In principle, short wavelengths can beobtained using high frequency modulated waves in a very turbid medium.However, the amplitude of the modulated wave decreases exponentiallywith the modulation frequency. Therefore, the best resolution, i.e., theshortest wavelength, is obtained using the highest frequency which stillgives a measurable signal. The diffusion process limits the penetrationdepth at any given modulation frequency, because of the exponentialdecrease of the wave's amplitude, and also decreases the velocity oflight propagation.

The above described diffusion wave approach treats amplitude modulatedlight waves in scattering media using the framework of wave optics. Thephoton intensity, calculated as superposition of different waves,constitutes a scalar field, propagating at a constant velocity. At anygiven modulation frequency, the wave optics phenomenology of scalarfields is valid. Therefore, in the frequency-domain, the measurement andanalysis of light diffusing in tissues from several sources will undergoconstructive and destructive interference. Furthermore, wave refractionoccurs at a boundary between two different tissues. It causes adeviation of the direction of propagation of the wave front, and thusthere is a change in the amplitude and phase shift of the propagationwave. The direction change is a function of the ratio of the effectiveindex of refraction in the two tissues. In diffusional wave optics, onthe other hand, the wave's amplitude is exponentially attenuated as thewave propagates in the scattering medium. This attenuation is inaddition to the exponential attenuation caused by finite absorption ofthe medium.

Amplitude modulated waves propagate coherently in the scattering medium;this is crucial for image reconstruction. It is possible to accuratelymeasure in real time, the average intensity, amplitude, and phase of thewave front over a large area using a single detector or an array ofdetectors applying well-established frequency-domain methods.

The emitters are varied sequentially in phase starting with the firstemitter in the line and followed by subsequent emitters. Each emitteremits a spherical wave and propagation of the resultant beam isperpendicular to the wavefront. If all the transmitter delays are equal,the beam travels straight ahead. Delay lines which produce variabletransmitter delays can be used to obtain appropriate phasing forsteering the beam across the tissue. The same principle can apply duringreception.

There are two important aspects of imaging as envisioned by the presentinvention. The first is a geometrical aspect and the second is phasingof the transmitters and receivers.

It is also possible to construct a two-dimensional array fortwo-dimensional pointing (e.g., FIG. 2C). The multiplexing switches usedwith these arrays can be constructed as an integral part of the arrayand can consist of field effect transistors arranged so that access toany element may be obtained by the application of two adverse signals.

In addition to electronic scanning, the two-dimensional scanning can beachieved by moving the array of sources and detectors in a regularpre-determined pattern in a plane parallel to that being investigated inthe subject. For maximum detection, the detector is places in the planeof the photon density gradient of the resulting field created by thearray of sources. The plane of the photon density gradient is swept asthe array moves. In this sweeping action, as a strongly or weaklyabsorbing object enters the radiation field, the detector registers afield imbalance due to the above described refraction of the propagatingradiation. A two-dimensional image is formed by storing the informationwhile the probe is moved across the subject. Several scans in differentimaging planes are envisioned by the invention. If the system isduplicated or time shared in two other faces of a cube, an algorithmwould be used to provide a 3-dimensional picture of the object bytriangulation. For a linear array of sources, there is a plane in whichthe null is sensitively detected, and the intersection of three planes(particularly at orthogonal intersection) defines the location of ahidden absorber. The data storage is accomplished electronically.

The detector detects the intensity and the phase shift of the radiationthat has migrated in the subject. The phase shift depends on the tissueproperties, i.e., absorption and scattering. For the low frequencies thephase shift is proportional to ((l-g)μ_(s)/μ_(a))^(½) and for the highfrequencies proportional to 1/μ_(a). To obtain desired penetrationdepth, appropriate frequency for both master oscillator 22 and localoscillator 26 is chosen; however, the phase relationship of the laserdiodes is maintained.

Different types of phased arrays are designed for optimal examinationand imaging of different human organs (e.g., human head or breast). Forexample, a mosaic of optical input ports and optical detection portsdefined by positions of optical fibers attached to a skull cap may beused. A standardized mapping may be developed also using x-raytechniques. Contrast labeling of different physiological structures willaid the visualization and orientation. The amplitude and phase of thesignals can be monitored on a precision oscilloscope. In order to scanthe phased array past a fixed object of approximately known position, asin needle localization procedures, the location of the input and outputports will be scanned past the object and the position of maximum phaseshift will be recorded in one-dimension; however, detection in two andthree dimension can be performed in the same way.

In the preferred mode of operation, the array of sources is phased 180°apart, as shown in FIG. 8A. There is a sharp 180° transition of photondensity wave, a large gradient in photon density, from S₂, S₂ sources tothe S₃, S₄ sources. Thus, the radiated field gives an amplitude null anda phase transition of 180° corresponding to the y-z plane, i.e.,perpendicular to the detector. If a larger number of similarly phasedsources is used, the transitions are even sharper. The array produces auniform photon density pattern on each side of the array, as shown inFIGS. 8B and 8C. If an absorbing object is placed in this directionalfield of diffusing optical waves, imbalance in the photon density ismeasured. The detection of a hidden object is accomplished bytranslating the experimental transmitter-receiver system of FIG. 8A.

In addition to the mechanical scanning achieved by moving of theinput-output port system, electronic scanning can be performed using themultiple source and multiple detector system of FIG. 2. As shown in FIG.2B for an array of five sources, there is a 180° phase transition in theresulting migrating field due to the 180° phase difference between theantiphased sources radiating amplitude modulated light. The plane of the180° phase transition can be shifted in parallel by appropriatelyvarying the apportionment of 0° and 180° phases on the sources. This isperformed by sequentially switching the phase of the sources by 180°. Ineach case, the detection port located on this plane is used forcollecting the data. As the sources are electronically switched by 180°,the detection array can be also electronically switched from onedetection port to another. The signal from the receiving optical fiberis coupled to one shared PMT detector. However, the system can alsoinclude several detectors. If the systems of FIGS. 1 or 1A are used, theelectronic source scanning can be combined with synchronous mechanicalmovement of the detection port.

In general, the invention utilizes the photon density gradient createdin the migrating field since it increases the resolution of thedetection. As known to one skilled in the art, the photon densitygradient formed by interference effects of introduced waves can becreated not only by appropriate phasing of the sources but also by othermethods such as appropriately spacing the sources, creating an imbalancein the radiated intensity of the individual sources, and other. Theimbalance may be achieved by modulating the amplitude of one source withrespect to another; this displaces the null in the correspondingdirection. Furthermore, the introduced signal can be encoded by thefrequency or a selected phase.

FIG. 8A shows the arrangement of the input ports 11 to 17 and detectionport 19 of FIG. 1. As described above, light of each laser diode 12through 18 is intensity modulated at the 200 MHz frequency. Wavelengthof the intensity modulated radiation is$\lambda = \left( \frac{4\pi \quad {c/n}}{3f\quad \mu_{s}} \right)^{\text{1/2}}$

wherein f is the modulation frequency of 200 MHz, μ_(g) is thescattering factor which is approximately 10 cm⁻¹ in an Intralipidsolution with refractive index n, and c is 3×10⁸ cm/s. Thus, theexpected wavelength is about 7 cm. The input ports S₁, S₂, S₃, and S₄are set 3.5 cm apart and are anti-phased by 180° using a push pulltransformer. The antiphased array creates a large gradient in photondensity chosen to take advantage of the destructive interference withthe null detection. The laser diodes emitting 754 nm light are intensitymodulated at 200 MHz using master oscillator 22, and the localoscillator 26 is operating at 200.025 MHz to perform the dynodemodulation of PMT detector 24. The detected intensities and phase shiftsof an x-direction scan (FIG. 8A) of detection port 19 are plotted inFIGS. 8B and 8C, respectively. As expected, the intensity has a sharpminimum in between sources S₂ and S₃ where the phase is changed 180°.The peak width at half maximum is about 2 cm. In addition to thex-direction scan of the detection port, the detection port was scannedin y-direction wherein, as expected, no variation was observed.

Referring to FIG. 9A, cylindrical objects of different diameter, d, werescanned using the previously described phased array. The objects wereplaced in the middle of the linear array displaced 2.5 cm from thex-axis. The detection port was located on the x-axis and each object wasmoved parallel to the x-axis at the 2.5 cm y displacement. The intensityand phase shift detected at different locations are plotted in FIGS. 9Band 9C, respectively. The intensity pattern for each moving object hastwo maximum and one minimum when the scanned object was located at x=0,y=2.5 point during its scan along the x-axis. At this point, a largephase change is detected, as shown in FIG. 9C. The phase detection hasinherently larger resolution of a localized absorber; a hidden object ofsize as small as 0.8 mm can be detected.

The response due to different absorption of the hidden object wasstudied using a 5 mm cylinder of different absorption coefficientscanned by the 4 element phased array of FIG. 9A. The detected phasechange is shown in FIG. 9D. The 5 mm black rod displays the largestphase change due to its high absorption, and the cylinder filled withcardiogreen 3.5 mg/l which has absorption coefficient μ_(a)=200 cm⁻¹shows the smallest phase change. In scanning of a hidden object, theseexperiments correspond to mechanically displacing the source detectorsystem, or electronically scanning the subject.

Scanning of two objects of a different diameter is shown in FIG. 10A.Two cylinders of different diameter are scanned across the four elementphased array located on the x-axis. The detection port in located at y=5cm. In FIG 10B the detected phase change in plotted against thedisplacement of these objects. Curve A represents the phase change oftwo cylinders of diameters 5 mm and 10 mm separated 3 cm apart. Curve Bwas measured using 16 mm cylinder instead the 5 mm cylinder. In thiscase, wherein the two cylinder separation is smaller, the phase detectorcan not resolve the two objects.

The imaging resolution is increased by increasing the number of elementsof the phased array, since the main lobe of the resultant beam becomesmuch sharper, the gradient of photon density is larger. Phased arrays ofdifferent number of elements and different shapes are used for imagingdifferent organs. For example, in tumor imaging, the four element phasedarray of FIG. 8A having an approximately linear shape can be used forimaging of the brain. On the other hand, a rectangular or a circularphased array would be used for imaging of a hidden tumor in the breast.The modulation frequency and the element spacing is adjusted to obtainproper focussing in each case.

In general, an imaging system will operate using the following modes ofoperation that arise from the above-described principles. In the firstmode of operation, a series of zero phased, appropriately spaced sourcescreate photon diffusion waves. One or more detectors sensitive to aselected wavelength detect the phase and the amplitude of the migratingwave. Individual sources and detectors may be coded and activatedaccording to selected detection and display schemes. The second mode ofoperation uses a series of sources phased at 0° and 180° (or any otheroffset phase that gives adequate sensitivity) with respect to eachother. The detector set at the null point of the array detects changesin the phase at the null point. Each detector may use an interferencefilter to limit its sensitivity to a selected wavelength. The third modeof operation may further complement the second mode by not onlydetecting the phase transition but also the amplitude null. The mostsensitive detection is achieved when a hidden object is located in themidline plane of the 0°-180° signal. An object is located using bothsignals and their appropriate integrals or derivatives are used toenhance the resolution of the system. The display will also utilizeinformation from several wavelengths, for example, when 750 nm and 850nm sources are used, the signal difference provides information aboutthe hemoglobin oxygenation and the sum about the blood concentration.Other wavelengths sensitive to endogenous or exogenous tissue pigmentsmay be used. The same source array may be designed to operate in allthree modes of operation. A computer supervisory system selects asuitable mode of operation for optimal sensitivity.

Referring to FIG. 11, a single wavelength localization system 83 employsa conical scanner 85 that introduces optical radiation of a selectedwavelength from four laser sources 87 to tissue 8. The relationship ofthe introduced patters is selected so that the resulting introducedradiation pattern forms a cone scanning in the examination space. Theoperation principles of array 87 were described in connection with FIGS.2A, 2B and 2C. Oscillator 62 generates a 200.025 MHz drive signal 91that is introduced to modulators 90 a and 90 b. Furthermore, the phaseof the drive signal is shifted by 90° in modulator 90 a relative to thephase of the drive signal is modulator 90 b, and the phase signals arevaried over time at 60 Hz. Each of the quadrature phase signals (92, 93)are splitted in splitter 89 a and 89 b to from an in-phase andanti-phase drive signals. The four drive signals drive four laser diodeslabeled N, S, W and E of array 87. Thus array 87 generates a scanningconical signal (88) that includes a sharp phase change in the center ofthe signal cone. Array 87 has four 780 nm laser diodes, but otherwavelengths selected for a high sensitivity to a tissue component may beemployed. Furthermore a multi-wavelength array can also be used.

The introduced diffusive photon density wave migrates in tissue 8 and isdetected at optical port 86 of an optical fiber connected to PMTdetector 75. As described above, the detected radiation is heterodynedusing a 200 MHz reference signal and the corresponding 25 kHz signal iscoupled to amplitude detector 96 and phase detector 79. Phase detector79 measures the phase shift between the introduced and detectedradiation patterns. The output of the phase detector is correlated withthe 60 Hz signals 92 and 93 to produce localization signalscorresponding to the N, S, W and E laser sources. The localizationsignals may be monitored using an oscilloscope.

When port 86 is symmetrically arranged in respect to the location of theradiation cone 88 and there is no field perturbation (i.e., no hiddenobject 9), the oscilloscope will display a circular pattern. In the samearrangement of cone 88 and port 86, if hidden object 9 is located in theradiation field, the oscilloscope pattern will no longer be symmetrical,e.g., the circular pattern may change to an elliptical pattern. Formaximum sensitivity, detection port 86 mechanically scans around tissue8 and is locked onto the scanning conical signal so that port 86 alwayspoints to the center of cone 88, i.e., port 86 is in the null location.

Referring to FIG. 11A, a phase modulation imaging system 100 includes atwo-dimensional phased array transmitter 102 connected to laser sources104. Electronics 120 drives laser sources 104 and also providesreference signals to the detection system. Optical detector 150 includesan optical input port 152 defined by a relatively large area opticalfiber 154 connected to a PMT detector 156.

Phased array transmitter 102 includes a horizontal array 106 and avertical array 112 of input ports connected by a set of optical fibers(not shown in FIG. 11) to laser sources 104 that include 754 nm and 816nm laser diodes labeled a and b, respectively. Diodes 107, 108, 109, and110 of the horizontal array 106 are driven by a push-pull transformer122, and diodes 103, 114, 115, and 116 of the vertical array 112 aredriven by a push-pull transformer 124. The resolution of the system maybe increased by adding more sources.

The horizontal sources are intensity modulated at a frequency ofapproximately 200.025 MHz generated by 200.025 MHz oscillator 124 and ahorizontal TV scan drive 128 generating a saw-tooth signal of 60 Hz. Ahorizontal reference signal 127 of 25 kHz supplied to phase detector 162is produced in a mixer 126 by mixing the 200.025 MHz signal fromoscillator 124 and a 200 MHz signal from oscillator 121. The verticalsources are intensity modulated at a frequency of approximately 200.2MHz generated by a 200.2 MHz oscillator 134 and a vertical TV scan drive138 generating a saw-tooth signal 139 of 1 kHz. A vertical referencesignal 137 of 200 kHz supplied to phase detector 164 is produced in amixer 136 by mixing the 200.2 MHz signal from oscillator 134 and the 200MHz from oscillator 121.

The emitted light of either 754 nm or 816 nm, alternated at 60 Hz by achopper, migrates in the examined tissue as described above and isdetected at input port 152. The detected light is heterodyned at PMTdetector 156 receiving a reference 200 MHz signal from oscillator 121.The detector signal is then filtered at 25 kHz and 200 kHz using filters158 and 160, respectively. Phase detectors 162 and 164, receiving 25 kHzand 200 kHz reference signals, respectively, determine at each frequencythe phase shift of the detected light in respect to the introducedlight.

As described above, the phase shift and the related optical pathlengthof the migrating photons directly reflect the tissue properties. System100 can distinguish the differences in the phase shift of the lightemitted from horizontal array 106 and vertical array 112 since theemitted light from each array is modulated at slightly differentfrequency.

Transmitter array 102 is designed to reflect the geometry of theexamined tissue and a possible location of hidden objects. The hiddenobjects A, B, and C of FIG. 11 targeted by array 102 are 3 to 4 cm inthe scattering medium. Thus, array 102 has the input ports spacedapproximately 1 cm apart and equidistantly from the center. Detectionport 152 is located about 7 to 10 cm from transmitter 102 and may bemechanically scanned in correlation with the total introduced field ofarray 102.

PMT detector 156 receives signals from the horizontal and verticalarrays. The modulation offset vertical frequency of the waveform isabout 10 times higher than for the horizontal waveform since therepeatability of the vertical scan is higher than the repeatability ofthe horizontal scan. Approximately the same frequency difference is usedfor the horizontal and vertical TV scans. The output from phasedetectors 162 or 164 represents the phase value as detected along thehorizontal axis and the vertical axis. A localized absorbing orscattering object (e.g., a tumor, localized bleeding) will cause a“resonance curve” type response. The detected phase shifts for eachsignal is differentiated (166 and 168) to “sharpen” the chances andincrease the resolution. The horizontal and vertical outputs are addedin a summing amplifier 170 and are coupled to a video input of a 500line TV display 180. The display may be graded in a gray scale or afalse color scale. The resolution achieved in the above describedone-dimensional experiments can be further improved and thesignal-to-noise ratio enhanced by employing a computer storage of thescanned data and integrating over a number of scans and using contrastenhancing algorithms. Alteratively, a “slow scan” TV may be used withnarrow banding of the outputs of the phase detectors.

System 100 may also include an amplitude detector 157 that detects theamplitude of the detected radiation at the 25 kHz and 200 kHzfrequencies. The detected amplitude signals are manipulated the same wayas the phase shift signals and fed to display 180. The use of both theamplitude signals and the phase signals improves resolution of theimage.

FIG. 11B shows diagrammatically a low frequency imaging system 190 thatemploys techniques similar to the ones used in system 100 of FIG. 11A. Asource array 192 emits diffused waves that propagate in tissue 195 andare detected by detectors 200. The highest resolution is achieved when ahidden object is located on the null line of the diffused waves. Thesystem operates at about 50 MHz, to use instead of the laser diodesLED's and instead of the PMT detectors Si diodes. Oscillators 202 and204 drive phase splitters 206 and 208, respectively, that provide twointensity modulated voltage signals shifted 180° with respect to eachother. The 0° and 180° signals drive 750 nm and 850 nm, LED sourceswhich are multiplexed by switches 210 and 212 to operate sources of onewavelength at the same time. The modulated diffuse waves are detected bythe Si diodes that include a wavelength specific interference filter,and the detector signal are converted from 50 MHz and 50.01 MHzfrequencies to 20 kHz frequencies using mixers 226 and 228,respectively. Phase detectors 230 and 232 operating at 20 kHz determinethe phase shift of the detected signals. Both the phase shift signalsand the amplitude signals are used to image the hidden absorber on adisplay unit 240.

A 2-dimensional transmitter and receiver arrays are shown in FIG. 12A.The spacing of the input ports can be varied depending on the frequencyof operation, expected location of the hidden object and the shape ofthe examined organ. FIG. 12B shows diagrammatically an imaging systemutilizing the 2-dimensional transmitter and receiver arrays 250 and 255that can be switched on electronically. A master oscillator 262 and alaser driver 260 drive a pair of in-phase and anti-phase laser diodes,e.g., the first and the third diode of Y-array and Z-array. A set ofelectronic switches is used to connect a different set of laser diodesevery 10 msec. A set of optical fibers transmits the detected light to aPMT detector 264 that also receives a reference 200.025 MHz signal froma local oscillator 266.

The heterodyned resulting signal is sent to a phase detector 272 thatmeasures the phase shift of the detected radiation. The measured phaseshift is further manipulated to enhance the detected changes on a CRTdisplay 276 which has the same 10 msec time base as electronic switches263. Differenciator 274 takes a derivative of the phase shift signal;this intensifies the crossover of the phase shift shown in FIGS. 8c, 9 cand 9 d.

Alternative Embodiments

In addition to the above described directional detection, the presentinvention envisions imaging systems constructed to calculate the averagemigration pathlengths. Referring to FIG. 4, in such system the drivesignal from oscillator 60 is introduced to a selected laser diode 64 a,. . . , 64 n or 66 a, . . . , 66 n using switches 61 a, . . . , 61 n.The intensity modulated radiation of each laser diode is coupled totissue 70 at an input port located at a precisely defined position. Adetection port located at another position detects radiation that hasmigrated in tissue 70. The detected signal is heterodyne mixed directlyat PMT detector 74. These signals are fed into the phase detectorwherein the phase and the intensity of the detected radiation aremeasured. The system may include several PMT detectors and phasedetectors (only one set of detectors is shown in FIG. 4) operatingsimultaneously or one detector scans the surface of tissue 70. The phaseshift and the intensity of the detected heterodyned signal depend on thetissue through which said scattered and absorbed radiation migrated.

The tissue properties are determined from the detected phase shift andintensity values and from the known input ports and detection portgeometries. The measured average pathlengths, <L>, can also bedetermined. The detected phase shift is converted to an effectivemigration pathlength <L> by using the low frequency approximationθ=2πf<L>n/c, wherein f is the modulation frequency, c is the speed oflight (3×10⁸ cm/s), and n is the refractive index of the medium.

To illustrate imaging by detecting migration pathlengths, we use anexample of photon migration in a tissue with a strongly absorbingobject, a perfect absorber(μ_(a)→∞) of radius R. Referring to FIGS. 5A,5B, and 5C the distribution of pathlengths defines an optical field thatexists between a point detector, D, and source, S, separated by distanceρ and located on the exterior of an examined tissue which is asemi-infinite, strongly scattering medium. As shown in FIG SA,infinitely far away from the field, a perfect absorber does not alterthe banana-shaped optical field of photons emitted by source S anddetected at detector D. As the object enters the optical field (FIG.5B), the photons which have migrated the farthest distance from D and Sare eliminated by the absorption process inside the perfect absorber ofradius R. Since photons which travel the longest pathlengths areabsorbed, the approach of an object shortens the distribution ofpathlengths, or alternatively, shortens the average pathlength <L>. Asthe object moves closer, and the optical field surrounds the object(FIG. 5C), some of the detected photons have travelled “around” theobject, which is detected as lengthening the distribution ofpathlengths. Thus, the average pathlength measurement can reveallocation of a strongly absorbing component of a tissue (e.g., tumor orlocalized bleeding).

Even though this pathlength computation approach requires in most casesextensive computational capabilities, it can yield useful information inthe localization procedures and can provide an useful supplement to theabove described directional approach.

What is claimed is:
 1. A spectroscopic system for examination ofbiological tissue comprising: multiple input ports arranged to introducelight at input locations into biological tissue and multiple detectionports arranged to collect light from detection locations of thebiological tissue, said multiple input and detection ports being placedat selected locations relative to a tissue region establishing for eachdetection port a plurality of average photon migration pathlengthsinside the tissue region between said input locations and said detectionlocations; at least one light source, operatively connected to aradiation pattern controller, constructed to generate light of awavelength in a range from visible to infrared, said light sourceoptically coupled to at least one of said input ports; at least onedetector, operatively connected to said radiation pattern controller,constructed and arranged to detect light of said wavelength that hasmigrated in the tissue region to at least one detection location andcorresponding at least one of said detection ports; said radiationpattern controller constructed to control patterns of said lightintroduced from said multiple input ports and constructed to controldetection of light migrating to said multiple detection ports to selecta null defined by introduced light patterns and said average photonmigration pathlengths; and a processor operatively connected to receivedetector signals from said detector and operatively connected to saidradiation pattern controller and arranged to process said detectorsignals relative to said patterns of said introduced light to determinea measured null in said patterns, said processor being arranged tocompare said null selected by said radiation pattern controller to saidmeasured null and arranged to examine the tissue region based on saidcomparison.
 2. The spectroscopic system of claim 1 wherein saidradiation pattern controller is arranged to select said light patternsin a way that light patterns emitted from two selected input portsdefine said null relative to one selected detection port, selection ofsaid input and detection ports corresponding to two selected averagephoton migration pathlengths between said selected input ports and saidselected detection port.
 3. The spectroscopic system of claim 2 whereinsaid radiation pattern controller is further arranged to directsequential introduction of said light from said selected input ports. 4.The spectroscopic system of claim 2 further comprising an imagerconnected to receive data from said processor and arranged to image saidtissue region based on said measured null for a plurality of said inputand detection locations.
 5. The spectroscopic system of claim 4 whereinsaid imager is arranged to receive from said processor said measurednull data sequentially detected for a plurality of input and detectionlocations arranged in a way that for each detection location there areat least two substantially symmetric input locations.
 6. Thespectroscopic system of claim 1 wherein at least two from said multipleinput ports provide at least two input locations arranged substantiallysymmetrically relative to at least one detection location of at leastone of said detection ports, and said radiation pattern controller beingarranged to select said light patterns in a way that light patternsintroduced at said two input locations define said null relative to saidsymmetrically arranged detection location.
 7. The spectroscopic systemof claim 6 wherein said radiation pattern controller being arranged tolocate said null at said symmetrically arranged detection location. 8.The spectroscopic system of claim 6 further including at least oneoscillator operatively connected to said at least one light source beingarranged to generate light modulated at a frequency of said oscillator,said at least one detector including a frequency filter arranged to passsaid detector signal of said frequency, and said radiation patterncontroller being arranged to control said introduction of said modulatedlight from said light source at said input locations and said detectionby said detector at said detection location.
 9. The spectroscopic systemof claim 8 wherein said oscillator operates at said frequency thatenables resolution of the phase shift that originates during migrationof photons in the tissue region.
 10. The spectroscopic system of claim 6including a plurality of oscillators operatively connected to aplurality of said sources arranged to generate light being modulated atfrequencies of said oscillators, and several said detectors includingfrequency filters arranged to pass said detector signals of saidmodulation frequencies, said radiation pattern controller being arrangedto control emission of said modulated light from said individual lightsources and detection of said modulated light by said light detectorsand said filtered detector signals from said frequency filters providedto said processor.
 11. The spectroscopic system of claim 10 wherein saidradiation pattern controller is further arranged to direct sequentiallysaid emission of said light modulated at different frequencies.
 12. Thespectroscopic system of claim 10 wherein said radiation patterncontroller is further arranged to direct simultaneously said emission ofsaid light modulated at different frequencies.
 13. The spectroscopicsystem of claim 10 further including at least one phase detectorconnected to receive said detector signal, said oscillators areconstructed to operate at said frequencies that enable resolution ofphase shifts that originate during photon migration in the tissueregion, and said phase detector being arranged to provide to saidprocessor said phase shifts corresponding to said average migrationpathlengths between said input and detection locations.
 14. Thespectroscopic system of claim 13 wherein said phase detector is furtherarranged to provide to said processor amplitudes of said detectorsignals corresponding to migration between said input and detectionlocations.
 15. The spectroscopic system of claim 13 wherein saidoscillators operate at said frequencies between 30 MHz to 200 MHz. 16.The spectroscopic system of claim 13 wherein said phase detector is aheterodyne phase detector.
 17. The spectroscopic system of claim 13wherein said phase detector is a homodyne phase detector.
 18. Thespectroscopic system of claim 13 wherein said processor is furtherarranged to calculate said average migration pathlengths between saidinput and detection locations.
 19. The spectroscopic system of claim 1further comprising an imager connected to receive data from saidprocessor and arranged to image said tissue region based on saidmeasured null for a plurality of said input and detection locations. 20.The spectroscopic system of claim 1 wherein said radiation patterncontroller is arranged to control detection by two of said detectors ina way that light pattern emitted from one of said input ports definessaid null relative to said two detection ports and corresponding twoaverage photon migration pathlengths between an input location of saidone input port and detection locations of said two detection ports. 21.The spectroscopic system of claim 20 wherein said radiation patterncontroller is further arranged to control sequential detection of saidlight at said detection ports.
 22. The spectroscopic system of claim 20further comprising an imager connected to receive data from saidprocessor and arranged to image said tissue region based on saidmeasured null for a plurality of said input and detection locations. 23.The spectroscopic system of claim 22 wherein said imager is arranged toreceive from said processor said measured null data sequentiallydetected for a plurality of input and detection locations arranged in away that for each detection location there are at least twosubstantially symmetric input locations.
 24. The spectroscopic system ofclaim 1 wherein at least one input location of at least one of saidinput ports is arranged substantially symmetrically relative to at leasttwo detection locations of at least two of said multiple detectionports, said radiation pattern controller being arranged to controldetection by two of said detectors in a way that the detection at saidtwo detection ports relative to the light pattern emitted from said atleast one input port defines said null corresponding two average photonmigration pathlengths between said at least one input location and saidat least two detection locations.
 25. The spectroscopic system of claim24 further including at least one oscillator operatively connected tosaid at least one light source being arranged to generate lightmodulated at a frequency of said oscillator, said at least two detectorsincluding at least two frequency filters arranged to pass said detectorsignals of said frequency, and said radiation pattern controller beingarranged to control said introduction of said modulated light from saidlight source at said input location and said detection by said twodetectors at said two detection locations.
 26. The spectroscopic systemof claim 25 wherein said radiation pattern controller is furtherarranged to direct sequentially said detection of said light modulatedat said frequency.
 27. The spectroscopic system of claim 25 wherein saidradiation pattern controller is further arranged to directsimultaneously said detection of said light modulated at said frequency.28. The spectroscopic system of claim 25 further including at least onephase detector connected to receive said detector signals from saiddetectors, said oscillator is constructed to operate at a frequency thatenables resolution of a phase shift that originates during photonmigration in the tissue region, and said phase detector being arrangedto provide to said processor said phase shift corresponding to saidaverage migration pathlengths between said input and detectionlocations.
 29. The spectroscopic system of claim 28 wherein said phasedetector is further arranged to provide to said processors amplitudes ofsaid detector signals corresponding to migration between said input anddetection locations.
 30. The spectroscopic system of claim 28 whereinsaid oscillator operates at said frequency between 30 MHz to 200 MHz.31. The spectroscopic system of claim 28 wherein said phase detector isa heterodyne phase detector.
 32. The spectroscopic system of claim 28wherein said phase detector is a homodyne phase detector.
 33. Thespectroscopic system of claim 28 wherein said processor is furtherarranged to calculate said average migration pathlengths between saidinput and detection locations.
 34. A spectroscopic method of examinationof biological tissue comprising: providing multiple input ports forintroducing light at input locations into biological tissue and multipledetection ports for collecting light from detection locations of thebiological tissue; positioning said multiple input and detection portsat selected locations relative to a tissue region to establish for eachdetection port a plurality of average photon migration pathlengthsinside the tissue region between said input locations and said detectionlocations; generating light having a wavelength in a range from visibleto infrared and optically coupling said generated light to said inputports; detecting, at least one of said detection ports, light of saidwavelength that has migrated in the tissue region to at least onedetection location; controlling patterns of said light to be introducedfrom said multiple input ports and controlling detection of lightmigrating to said multiple detection ports to select a null defined bysaid introduced patterns and said average photon migration pathlengths;receiving detection signals and processing said detection signalsrelative to said patterns of said introduced light to determine ameasured null in said patterns; and comparing said selected null to saidmeasured null to examine the tissue region based on said comparison. 35.The spectroscopic method of claim 34 wherein said controlling includesselecting said light patterns in a way that light patterns emitted fromtwo selected input ports define said null relative to one selecteddetection port corresponding two selected average photon migrationpathlengths between said selected input ports and said selecteddetection port.
 36. The spectroscopic method of claim 35 wherein saidcontrolling further includes directing sequential introduction of saidlight from said selected input ports.
 37. The spectroscopic method ofclaim 34 wherein said positioning includes providing at least two inputlocations arranged substantially symmetrically relative to at least onedetection location, and said controlling includes selecting said lightpatterns in a way that light patterns introduced at said two inputlocations define said null relative to said symmetrically arrangeddetection location.
 38. The spectroscopic method of claim 37 whereinsaid selecting includes locating said null at said symmetricallyarranged detection location.
 39. The spectroscopic method of claim 37further comprising imaging said tissue region based on said measurednull for a plurality of said input and detection locations.
 40. Thespectroscopic method of claim 34 wherein said generating includesmodulating at a selected frequency said light, said detecting includesfiltering said detection signal of said frequency, and said controllingincludes controlling said introduction of said modulated light from saidlight source at said input locations and said detection by said detectorat said detection location.
 41. The spectroscopic method of claim 40wherein said frequency that enables resolution of the phase shift thatoriginates during migration of photons in the tissue region.
 42. Thespectroscopic method of claim 41 further including phase detecting saiddetection signal relative to said modulated light to measure phaseshifts corresponding to said average migration pathlengths between saidinput and detection locations.
 43. The spectroscopic method of claim 42further including detecting amplitudes of said detector signalscorresponding to migration between said input and detection locations.44. The spectroscopic method of claim 41 wherein said phase detectingincludes homodyne detection.
 45. The spectroscopic method of claim 41wherein said phase detecting includes heterodyne detection.